Failure of hemodialysis vascular access and other vascular grafts becomes evident as compromise of the lumen of the native vessel (vein or artery) or of the prosthetic conduit at o or away from the anastamotic site. Compromise of the lumen manifests as either stenosis or occlusion and is a result of either intraluminal thrombus and/or a vasculoproliferative response. The etiology of graft failures may be related to a variety of physical (e.g., shear stress causing hemodynamic disturbance), chemical and/or biological stimuli as well as infection and foreign body rejection which may explain why fistulae which do not involve a foreign body (in this case, for example, polytetrafluroethylene, PTFE) remain patent for a longer time compared to vascular access grafts that involve interposition of a PTFE graft.
The present invention relates generally to therapeutic implant, apparatus and methods useful for preventing, suppressing (inhibiting) or treating failure of hemodialysis vascular access and other vascular grafts.
Vascular access grafts, specifically, hemodialysis access grafts are well known to the art. Approximately 100,000 vascular access procedures are performed yearly in the United States. Hemodialysis vascular access can be constructed in one of several ways: as an arterio-venous fistula (e.g.; Brecisa-Cimino), or as a graft, interposing either prosthetic (e.g., PTFE) or biologic tissue (e.g., vein) between the artery and the vein. Such grafts are usually constructed using a tubular or cylindrical segment of suitably bio-compatible, substantially inert material such as polytetrafluoroethylene (PTFE). In fact, PTFE is the most common material used for prosthetic dialysis access. In one approach, a segment of PTFE is surgically interposed between an artery and a vein in the arm, forearm or thigh. The graft is then available for repeated vascular access for performing hemodialysis.
Subsequent to placement of the access graft the sutured sites in the artery and the vein undergo healing. Sixty percent of these grafts fail each year, usually because of narrowing (stenosis) at the venous end. Similar lesions develop in PTFE grafts placed in the arterial circulation, where there is a similar tendency for the distal end of the graft to be affected. Dysfunction or failure of veing grafts and/or other graft conduits used in coronary artery bypass graft surgery or in peripheral vascular surgery (e.g., aorta-iliac, femoral-femoral, femoral-popliteal, femoral tibial, etc.) are well known. Development of arterial access graft stenosis is not as rapid as development of access graft stenosis at the venous end. Proliferation and migration of smooth muscle cells resulting in intimal hyperplasia in the vein and the adjacent graft orifice has been described in human dialysis access stenosis. As the stenosis in the graft becomes progressively more severe, the graft becomes dysfunctional and hemodialysis is suboptimal. If the stenosis in the graft is not treated, it eventually leads to occlusion and to graft failure.
The reasons why the venous ends of access graft have such a marked propensity for narrowing are multifactorial. Features unique to this location include exposure to arterial pressures and arterial flow rates, dissipation of acoustic (vibratory) energy in the vessel wall and surrounding tissue, repeated puncture of the graft, and infusion of processed blood. In addition, the venous end of the graft may be bathed in mitogens released during passage of the blood through the dialysis tubing or during activation of platelets at the site of needle puncture.
Tissue samples collected from the graft-vein anastomosis site of stenotic PTFE grafts during surgical revision showed significant narrowing of the lumen and were characterized by the (i) presence of smooth muscle cells, (ii) accumulation of extra-cellular matrix, (iii) angiogenesis within the neointima and adventitia, and (iv) presence of an active macrophage cell layer lining the PTFE graft material. A large variety of cytokines and cell growth stimulating factors like platelet-derived growth factor (PDGF), basic fibroblast growth factor (bFGF), and vascular endothelial growth factor (VEGF) were expressed by smooth muscle cells/myofibroblasts within the venous neointima, by macrophages lining both sides of the PTFE graft, and by vessels within the neointima and adventitia. It has been suggested that macrophages, specific cytokines (bFGF, PDGF, and VEGF), and angiogenesis within the neointima and adventitia are likely to contribute to the pathogenesis of venous neointimal hyperplasia (VNH) a manifestation of the vasculoproliferative response in PTFE dialysis grafts.
Survival of patients with chronic renal failure depends on optimal regular performance of dialysis. If this is not possible (for example as a result of vascular access dysfunction or failure), it leads to rapid clinical deterioration and unless the situation is remedied, these patients will die. Vascular access dysfunction is the most important cause of morbidity and hospitalization in the hemodialysis population in the United States at an estimated cost of approximately one billion US dollars per annum. Venous neointimal hyperplasia characterized by stenosis and subsequent thrombosis accounts for the overwhelming majority of pathology resulting in PTFE dialysis graft failure. Despite the magnitude of the problem and the enormity of the cost, there are currently no effective therapies for the prevention or treatment of venous neointimal hyperplasia in PTFE dialysis grafts. Consequently, interventions aimed at the specific mediators and processes may be successful in reducing the very significant human and economic costs of vascular access dysfunction.
Once the stenosis has occurred, one of the current methods of treatment involves reduction or obliteration of the narrowing and restoration of blood flow through the graft (permitting the performance of adequate hemodialysis) by means of non-surgical, percutaneous catheter based treatments such as balloon angioplasty. Balloon angioplasty, in one aspect, involves deployment of a balloon catheter at the site of the blockage and inflating the balloon to increase the minimum luminal diameter (MLD) of the vessel by compressing the material causing the restriction against the interior of the vessel wall, thereby dilating the vessel. Depending upon the length and severity of the restriction, the procedure may be repeated several times (by inflating and deflating the balloon). When completed, the balloon catheter is withdrawn from the system.
Although balloon angioplasty can be used as a “stand alone” procedure, it is frequently accompanied by deployment of what is called a stent. A stent is an expandable scaffolding or support device which is placed within the vasculature to prevent mechanical recoil and reduce the chance of renarrowing (restenosis) at the site of the original restriction. Stents are either “balloon-expandable” or “self-expanding” and when deployed endovascularly, abut against the inner vessel wall. Whether or not a stent is placed, this form of treatment has a high risk of failure i.e., the risk of renarrowing (restenosis) at the treatment site is very high. Unless stenosis within the access graft can be effectively and permanently treated, graft failure tends to follow. In the event of graft failure, the patient has to undergo an endovascular procedure i.e., a non-surgical, catheter-based percutaneous procedure, repeat vascular surgery e.g., thrombectomy to “declot” the graft or to place another vascular access graft or a shunt (as it is sometimes referred to) at a different site, unless the patient receives a kidney transplant. Given the obvious problems of repeat surgery(ies) and the limited availability of transplants, there is a need for a treatment that is both effective and long lasting (durable) in the prevention and treatment of dialysis graft stenosis.
The vast majority of current approaches for reducing or preventing the vasculoproliferative response (believed to be the pathophysiological basis of restenosis), are based on treatment options that originate from within the vascular or graft lumen. One current, novel approach utilizes drug coated or drug impregnated stents which are then deployed within the lumen of the blood vessel. Examples of drugs used to coat stents include Rapamycin commercially available from the Wyeth Ayerst company (Sirolimus®), and Paclitaxel commercially available from the Bristol-Myers Squibb Company (Taxol®). In this stent-based approach, Rapamycin or Paclitaxel is gradually eluted from the stent and diffuses into the vessel wall from the intima (the innermost layer of the vessel wall) to the adventitia (the outermost layer of the vessel wall). Studies have shown that Rapamycin and Paclitaxel tend to inhibit smooth muscle cell proliferation.
Delivery from the perivascular or extravascular space through the arterial or vascular wall utilizing a synthetic matrix material (ethylene-vinyl acetate copolymer, EVA) together with an anticoagulant that also has antiproliferative properties e.g., heparin, has been suggested. There are two disadvantages of this approach: heparin is a soluble substance and rapidly disappears from the vascular wall and, ethylene-vinyl acetate copolymer is not biodegradable potentially raising concerns about long term effects, in vivo.
If a therapeutic agent is delivered locally using a matrix material-based system, the matrix material should preferably have the following characteristics:
1. The matrix material has to permit the loading of adequate quantity of the therapeutic agent.
2. The matrix material must elute the therapeutic agent at an appropriate, well defined rate.
3. The matrix material should preferably be implantable and biodegradable. Thus, physical removal of the matrix material from recipient's tissue following drug delivery would not be necessary and obviates concerns about the long term effects of the residual matrix.
4. Neither matrix material nor its biodegradation products should provoke a significant inflammatory or proliferative tissue response, nor should they alter or interfere with the recipient's natural defense systems or healing.
5. The device (comprising the matrix material and the drug) should be flexible enough to mould to the contours of the vasculature and
6. The device should be amenable to be fixed in place preventing its migration to an unintended location.
Polymer matrix materials used for drug delivery within the context of implantable devices can be either natural or synthetic. Examples include but are not limited to polymers composed of chemical substances like polyglycolic acid or polyhydroxybutyrate, EVA or natural polymers like collagen, fibrin or polysaccharides like chitosan. However, not all of these matrix materials are ideal; inappropriate features include poor mechanical characteristics, potential immunogenicity, and cost. In addition, some may produce toxic degradation products and induce inflammatory reactions or a proliferative response.
A well known biocompatible, biodegradable, resorbable matrix material for drug delivery is collagen. The use of collagen as a material for fabrication of biodegradable medical devices is and has undergone serious scrutiny. U.S. Pat. Nos. 6,323,184, 6,206,931; 4,164,559; 4,409,332; 6,162,247. One current focus involves delivery of pharmaceutical agents including antibiotics and physiologically active proteins and peptides such as growth factors.
Under scanning electron microscopy, the collagen matrix has a morphology of condensed laminated film with a textured surface and a range of pore sizes. It can be produced in a wide range of effective pore sizes from 0.001 microns to 100 microns or even larger. This internal pore network (porous material) creates a high surface area and serves as a microreservoir for storage and delivery of the therapeutic agent. Several features make collagen an excellent and ideal matrix material for drug delivery. Collagen exhibits a high degree of flexibility and mechanical durability, as well as intrinsic water wettability, semipermeability and consistent flow characteristics. More importantly, collagen, a naturally occurring substance is biodegradable and non-toxic. In addition, collagen has favorable biodegradation characteristics and time to complete degradation or resorption i.e., durability of the collagen matrix for drug delivery can be modified.
A second protein matrix suitable for drug delivery is fibrin. A fibrin matrix is comprised of cross-linked fibrin units that are a reticular network of thrombin-modified fibrinogen molecules. This matrix is similar to a natural blood clot. In contrast to natural clot, the size of pores in a fibrin matrix can be controlled and varies from 0.001 millimicrons to 0.004 millimicrons, so-called micropores. The differences in pore sizes between collagen and fibrin matrices permit the binding of therapeutic agents with distinct rates of drug release. The ability to control bleeding, to remain firmly fixed in place, and to be naturally biodegradable have all made fibrin a good matrix material for drug delivery and confers fibrin some advantages over synthetic matrices. Most of the early applications of fibrin as a matrix were for delivery of antibiotics and other biologics.
The fibrin matrices are prepared in a dry granular form. (cf., PCT/EP99/08128). This formulation, manufactured by HyQ Solvelopment, Buhlmhle, Germany, contains D-mannitol, D-Sorbit, fibrinogen-aqueous solution, and a thrombin-organic suspension. The formulation is manufactured by fluid bed granulation. The applications for dry fibrin are manifold: wound closure, promotion of healing, and homeostasis. However, application for drug delivery is limited since such a formulation does not allow for a target-oriented shaping of solid particles around the vessel wall and delivery of exact dosages is difficult. Porosity and capacity of dry fibrin particles are low, physical stability is poor.
Another group of potentially useful resorbable, natural polymer matrix material is chitosan. Chitosan has proven to be a useful biocompatible aminopolysaccharide and a matrix for controlled release of the agent for local delivery. Chitosan implants cause no systemic and local side effects or immunologic responses, and are suitably biodegradable. Chitosan can be prepared from the degradation of slow chitin (molecular weight 1×106) using high temperature sodium hydroxide hydrolysis to a molecular weight of 5×105. The inability to control porosity is a disadvantage of this matrix material.